The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to the acquisition of an NMR image data set using a fast pulse sequence.
Any nucleus which possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant .gamma. of the nucleus). Nuclei which exhibit this phenomena are referred to herein as "spins".
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment M.sub.z is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, M.sub.z, may be rotated, or "tipped", into the x-y plane to produce a net transverse magnetic moment M.sub.t, which is rotating, or spinning, in the x-y plane at the Larmor frequency.
The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation signal B.sub.1 is terminated. In simple systems the excited spins induce an oscillating sine wave signal in a receiving coil. The frequency of this signal is the Larmor frequency, and its initial amplitude, A.sub.0, is determined by the magnitude of the transverse magnetic moment M.sub.t. The amplitude, A, of the emission signal decays in an exponential fashion with time, t: EQU A=A.sub.0 e.sup.-t/T*.sub.2
The decay constant 1/T*.sub.2 depends on the homogeneity of the magnetic field and on T.sub.2, which is referred to as the "spin-spin relaxation" constant, or the "transverse relaxation" constant. The T.sub.2 constant is inversely proportional to the exponential rate at which the aligned precession of the spins would dephase after removal of the excitation signal B.sub.1 in a perfectly homogeneous field. The practical value of the T.sub.2 constant is that tissues have different T.sub.2 values and this can be exploited as a means of enhancing the contrast between such tissues.
Another important factor which contributes to the amplitude A of the NMR signal is referred to as the spin-lattice relaxation process which is characterized by the time constant T.sub.1. It describes the recovery of the net magnetic moment M to its equilibrium value along the axis of magnetic polarization (z). The T.sub.1 time constant is longer than T.sub.2, much longer in most substances of medical interest. As with the T.sub.2 constant, the difference in T.sub.1 between tissues can be exploited to provide image contrast.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (G.sub.x, G.sub.y, and G.sub.z) which have the same direction as the polarizing field B.sub.0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.
Most NMR scans currently used to produce medical images require many minutes to acquire the necessary data. The reduction of this scan time is an important consideration, since reduced scan time increases patient throughput, improves patient comfort, and improves image quality by reducing motion artifacts. There is a class of pulse sequences which have a very short repetition time (TR) and result in complete scans which can be conducted in seconds rather than minutes. Whereas the more conventional pulse sequences have repetition times TR which are much greater than the spin-spin relaxation constant T.sub.2 so that the transverse magnetization has time to relax between the phase coherent excitation pulses in successive sequences, the fast pulse sequences have a repetition time TR which is less than T.sub.2 and which drives the transverse magnetization into a steady-state of equilibrium. Such techniques are referred to as steady-state free precession (SSFP) techniques and they are characterized by a cyclic pattern of transverse magnetization in which the resulting NMR signal refocuses at each RF excitation pulse to produce an echo signal. This echo signal includes a first part S+ that is produced after each RF excitation pulse and a second part S- which forms just prior to the RF excitation pulse.
There are two well known SSFP pulse sequences used to produce images. The first is called gradient refocused acquired steady-state (GRASS) and it utilizes a readout gradient G.sub.x to shift the peak in the S+ signal that is produced after each RF excitation pulse toward the center of the pulse sequence. This pulse sequence is shown in FIG. 3, where the NMR signal is an S+ gradient echo that is induced by the readout gradient G.sub.x. In two-dimensional imaging, a slice selection gradient pulse is produced by the gradient G.sub.z and is immediately refocused in the well-known manner. A phase encoding gradient pulse G.sub.y is produced shortly thereafter to position encode the acquired NMR data, and to preserve the steady-state equilibrium, the effects of the phase encoding gradient pulse are nullified by a corresponding G.sub.y rewinder gradient pulse after the NMR signal has been acquired and before the next pulse sequence begins as described in U.S. Pat. No. 4,665,365.
The second well known SSFP pulse sequence is called contrast enhanced fast imaging (SSFP-ECHO) and it utilizes the S- signal that is produced just prior to each RF excitation pulse. This pulse sequence is shown in FIG. 4, where the acquired NMR signal is an S- echo signal caused by the gradient refocusing of the transverse magnetization which would otherwise refocus at the next RF excitation pulse. The readout gradient G.sub.x is substantially different in this pulse sequence and includes a positive pulse prior to the actual readout pulse and a negative pulse after the readout pulse. The former pulse dephases the FID signal (S+) which might otherwise be produced during the data acquisition window, and the latter pulse causes the transverse magnetization to rephase during the next pulse sequence to produce the echo signal S-. For a more detailed discussion of the SSFP-ECHO pulse sequence, reference is made to an article by R. C. Hawkes and S. Patz entitled "Rapid Fourier Imaging Using Steady-State Free Precision", published in Magnetic Resonance in Medicine 4, pp. 9-23 (1987).
Because SSFP sequences employ RF excitation pulses with small tip angles and the magnetization is not allowed to recover after each pulse sequence, the image contrast due to spin density is not nearly as good as with conventional pulse sequences. Consequently, other image contrast enhancement methods have been proposed which rely on the different T.sub.1 and T.sub.2 constants of tissues. As described by A. Haase in "Snapshot Flash MRI Applications to T1, T2, and Chemical-Shift Imaging," Magnetic Resonance In Medicine, 13, 77-89 (1990), and D. Matthaei et al in "Fast Inversion Recovery T.sub.1 Contrast and Chemical Shift Contrast In High-Resolution Snapshot Flash MR Images," Magnetic Resonance Imaging, Vol 10, pp. 1-6, 1992, a series of SSFP pulse sequences may be preceded by one or more preparatory RF pulses which condition the spin magnetization to provide T.sub.1 or T.sub.2 enhanced contrast images. In some 2D acquisitions a single preparatory period followed by the acquisition of all the image views will suffice. However, in 3D acquisitions and in multi-slice 2D acquisitions the scan is divided into a series of data acquisition periods which are each preceded by a contrast enhancement preparatory period. In addition, and as described by J. P. Mugler et al in "Three-Dimensional Magnetization-Prepared Rapid Gradient-Echo Imaging (3D MP RAGE)," Magnetic Resonance In Medicine 15, 152-157 (1990); by M. Brant-Zawadzki in "MP RAGE: A Three-Dimensional, T1-Weighted, Gradient-Echo Sequence--Initial Experience in the Brain," Radiology 1992; 182: 769-775; and by J. P. Mugler et al in "T2-Weighted Three-Dimensional MP-RAGE MR Imaging," JMRI 1991:1:731-737; each data acquisition period is followed by a recovery period. Each recovery period enables the spin magnetization to return to thermal equilibrium from the lower steady-state equilibrium attained during the previous SSFP acquisition period. For example, in a typical 3D, T1-weighted, magnetization prepared scan, each prepatory period may require from 300 to 500 milliseconds, followed by n SSFP acquisition periods of between 6 to 12 milliseconds each, where n is the number of sections prescribed. This is followed by a recovery period which may require 1 to 2 seconds. This sequence might be repeated m times, where m is the in-plane resolution in the phase encoding direction. Typically, a 3D T1-weighted scan requires 64 sections and has an in-plane resolution in the phase encoding direction 256, for a total scan time of 548 seconds (9.13 minutes).
Rapid acquisition of fat suppressed images is essential in the diagnosis of breast lesions. SSFP pulse sequences inherently provide images with good T.sub.1 contrast. By injecting the patient with a contrast agent such as Gadolinium, the T.sub.1 constant of tumor cells is shortened relative to the surrounding muscle and fiber tissues and produces a characteristic pattern in an SSFP image. However, fat tissues also have a short T.sub.1 time, and they will also appear brightly in the image. These images must be acquired within two to three minutes in order to maximize the effect of the Gadolinium tumor contrast agent, and the signals from fat tissues must be suppressed so that the tumor can be seen against the background of other tissues which have a longer T.sub.1 time. In addition, these images must be rapidly acquired before the Gadolinium has perfused through normal tissues which tend to mask the signal from tumors.